When indicating and assessing the state of the art in the following description, appropriate citations in the generally accessible literature are cited:    [1] R. Bader et al., “A wide frequency range antenna system for multi-nuclear MRS and MRI.”, SMRM, 6th Annual Meeting, New York 1987, p. 406;    [2] P. K. Roschmann, “Surface coil for high-frequency magnetic fields for magnetic resonance examinations.”, U.S. Pat. No. 4,775,837;    [3] J. T. Vaughan et al., “High-frequency surface coils for clinical NMR imaging and spectroscopy.”, SMRM 12th Ann. Meeting, New York 1993, p. 1332;    [4] C. E. Hayes et al., “An efficient, highly homogenous radiofrequency coil for whole-body NMR-imaging at 1,5-T.”, J. Magn. Reson. 1985, 63: 622-628;    [5] W. A. Edelstein et al., “The relative sensitivity of surface coils to deep lying tissues.” SMRM, 4th Annual Meeting, San Francisco 1985, pp. 964-965;    [6] Silva and Merkle, “Hardware considerations for functional magnetic resonance imaging.” Concepts Magn. Reson. A, 2003; 16: 35-49;    [7] T. Mildner et al., “Functional perfusion imaging using continuous arterial spin labeling with separate labeling and imaging coils at 3 T.” Magn. Reson. Med. 2003, 49: 791-795;    [8] J. T. Vaughan et al., “High-frequency volume coils for clinical NMR imaging and spectroscopy.”, Magn. Reson. Med. 1994, 32: 206-218;    [9] J. T. Vaughan, “Radiofrequency volume coil.”, U.S. Pat. No. 5,557,247, 1996;    [10] J. T. Vaughan, “HF-coil for imaging system and use therein.”, U.S. Pat. 60/135,269;    [11] G. Bogdanov and R. Ludwig, “Coupled microstrip line transverse electromagnetic resonator module for high-field magnetic resonance.”, Magn. Reson. Med. 2002, 47: 579-593;    [12] P. Bottomley, “MRI tunable antenna and system.”, WO 03/0582283;    [13] R. F. Lee et al., “Lumped-element planar strip array (LPSA).”, Magn. Reson. Med. 2004, 51: 172-183;    [14] R. F. Lee et al., “Planar strip array (PSA) for MRI.”, Magn. Reson. Med. 2001; 45: 673-683;    [15] W. Driesel and H. Merkle, “3-Tesla-helmet coil customization to allow for more general fMRI studies.” MAGMA 2003, 16 (Suppl. 7); 163;    [16] X. Zhang et al., “Microstrip RF surface coil design for extremely high-field MRI and spectroscopy.” Magn. Reson. Med. 2001; 46: 443-450;    [17] J. D. Jackson, “Klassische Elektrodynamik.” W. de Gruyter, Berlin 1983;    [18] I. J. Bahl and D. K. Trivedl, “A designer's guide to microstrip line.” Microwaves, May 1997, pp. 174-179;    [19] F. Schnieder and W. Heinrich, “Model of thin-film microstrip line for circuit design.” IEEE Trans. Microwave Theory Tech. 2001, 49: 104-110;    [20] T. Itoh, “Overview of quasi-planar transmission lines.” IEEE Trans. Microwave Theory Tech. 1989, 37: 275-280;    [21] G. Zaharchuk et al., “Multislice perfusion and perfusion territory imaging in humans with separate label and image coils.” Magn. Reson. Med. 1999; 41: 1093-1098;    [22] R. Trampel et al., “Continuous arterial spin labeling using a local magnetic field gradient coil.” Magn. Reson. Med. 2002; 48: 543-546.    [23] M. D. Schnall, “Probes tuned to multiple frequencies for in-vivo NMR.” In: NMR Basic Principles and progress, 26: 33-63, Springer, Heidelberg, 1992.
These citations are designated in brackets [ ] in the text of the description by indicating the above reference numerals.
Since the arrangements for exciting and/or detecting magnetic resonance in the beginning of magnetic resonance technology had the shape of conductive loops wound similar to coils, they are still often designated today as coils, even if their actual design has no similarity with what is generally understood by a coil. In the text presented here the more comprehensive designation “antenna” is preferred, which in the meantime has also become customary in parts of the professional literature.
The HF field in the close range, that is, ranges of the examined object in the direct surroundings of the antenna (distances up to one wavelength) is of particular interest as regards the design of the HF antenna. The so-called far field should be avoided to the extent possible since it only results in undesired interactions with the surroundings. There is a broad spectrum for using such antennas in the medical area, e.g., when investigating the human brain. However, the invention can also be used with advantage in other areas of MRI and MRS and for other objects.
In MRI or MRS the object to be examined is brought into a strong static magnetic field (B0 field). If the object to be examined contains atomic nuclei with a nuclear spin different from zero the spins and the magnetic nuclear moments associated with them are oriented in the B0 field and the splitting of the nuclear spin energy level that degenerated in the field-free space takes place. The nuclear moments execute a precessional movement around the direction of the static magnetic field Bo. The frequency of the precession (Larmor frequency ω) is proportional to the B0 field strength:Ω=γ*B0.
The symbol “*” designates here and in the following the operator for multiplication. The proportionality factor γ is the gyromagnetic ratio of the atomic nucleus examined. The Larmor frequency is in a range of approximately 1 MHz to 1 GHz as a function of the B0 field strength. A macroscopic magnetization that is directed parallel to the direction of the static field B0 in thermal equilibrium results as the vectorial sum, normalized to the volume, of the magnetic nuclear moments in the examined object. In order to deflect the magnetization, that is, in order to generate a measurable magnetic resonance signal, electromagnetic HF power is radiated in pulses or also continuously into the examined object with the Larmor frequency ω by means of an HF transmitter and an HF antenna. When receiving, the voltage induced in the antenna is measured as response signal of the excited spins. Conclusions can be drawn about the physical and chemical structure of the object by evaluating of this “magnetic resonance signal”. In this manner morphological, metabolic and functional examinations are possible in biomedical applications. If a B0 field gradient is superposed on the homogenous B0 field the Larmor frequency becomes location-dependent. This quality is utilized in MRI for location coding (site coding).
The desired interaction between the spin of the examined object and the HF antenna results only for those locations in the examined object where the magnetic alternating field has a component orthogonal to the z direction, that is, orthogonal to the direction of the B0 field. This component is designated as the “B1 field”.
The customary transmitting/receiving antennas for magnetic resonance can be roughly subdivided into surface- and volume antennas. The surface antennas also include, among other things and in addition to simple conductive loops, loop-shaped structures of waveguides, see [1], [2], [3].
Typical representatives of volume antenna are, among others, saddle coils, solenoids and birdcage resonators. Combinations of surface antennas and volume antennas are also known. Volume antennas that generate a very homogenous HF field (B1 field) are in particular the birdcage resonators, see [4].
All these antenna types have disadvantages in their application. Thus, a very homogenous illumination of the object can be achieved with a volume antenna; however, a lesser sensitivity as regards the magnetic resonance signal is gained. In contrast thereto, surface antennas have a very high sensitivity to the magnetic resonance signal in areas close to the surface, that is, at a slight depth vertically to the antenna plane, see [5]. However, the illumination takes place very inhomogeneously and decreases rapidly with increasing distance. Surface antennas can be advantageously used for receiving the magnetic resonance signal approximately down to a depth corresponding to the diameter of the antenna.
All efforts to obtain an optimal antenna design are based on the desire to create an HF antenna that is as sensitive as possible only in the area of interest (therefore, in the “measuring volume”), that is, that excites the nuclei in the examined object only there and receives magnetic resonance signals only from there. It would be ideal if the illumination is as homogenous as possible in this range and the antenna has a high sensitivity for the actual measuring volume. These two requirements are mutually exclusive. As is known, in the direct vicinity of conductors the magnetic field is very inhomogeneous but the receiving efficiency S regarding the magnetic resonance signal is high:S˜B(r)/I,in which B(r) designates the magnetic field at distance r from the conductor and I designates the current through the conductor. The symbol “˜” designates “proportional to” and the symbol “/” designates here the operator for division. Inverse conditions are found at a greater distance from the conductors.
This is a problem if different locations in the measuring volume have different distances from the antenna. Thus, the design of HF antennas always represents a compromise between the two competing viewpoints. It is apparent from this that HF antennas must be optimized for corresponding applications.
A known measure to this end is to adapt the antenna in its form to the object to be examined in such a manner that the volume area of interest of the object is detected as well as possible by the magnetic field in the close range. Thus, anatomically adapted antennas were developed for investigating the human brain that can surround a human head like a helmet, see [6]. However, such a “helmet antenna” has certain disadvantages in its known embodiment, as will be explained in the following using a figure of the attached drawings.
The antenna according to FIG. 1 contains, in its simplest shown form, a ring-shaped conductive base R subdivided into four areas by four condensers CR distributed uniformly over its circumference. Halfway between every two adjacent ring condensers CR one of four conductive holders X1, X2, Y1, Y2 offset by 90° empties in, which are also subdivided for their part by condensers CB for tuning purposes and to avoid phase shifts. The two diametrically opposite holders X1 and X2 are connected at the other end by a first feed connection x1-x2. The other two diametrically opposite holder halves Y1 and Y2 are connected by a second feed connection y1-y2. Each pair of two diametrically opposite holders forms a circuit that is closed via base ring R.
Base ring R, the first pair of holders X1, X2 and the second pair of holders Y1, Y2 act as three partial antennas. Since these three partial antennas are arranged in three planes vertical to each other a decoupling between the three partial antennas can be achieved given an exactly symmetrical construction of the helmet antenna relative to the z axis. Only the two X and Y pairs of holders substantially generate a magnetic field vertically to the z axis, that is, an active B1 field, and are therefore active as concerns the magnetic resonance. Furthermore, a strong gradient of the B1 field, measured in the B0 direction (z direction) results in the case of the described helmet antenna because on account of the geometric form of the pairs of holders this field is much weaker in the area of the antenna base, where the holders are located further apart from each other, than in the area of the antenna top. This has the result that examinations with a usable signal-to-noise ratio (SNR) are possible only in the upper area of the brain.
In addition to this disadvantage there is yet a further problem that is based in the described helmet antenna concept. As was explained above, these antenna concepts use a ring-shaped base for the current flow between the two ends of each pair of holders far from the top. This results in a serious disadvantage that should be explained for one pair of holders. The explanation also logically applies to the pair of holders offset by 90°. The current penetrating from the one holder into the ring-shaped base flows half and half into the two parts of the ring-shaped base further, where it generates an additional magnetic field. This magnetic field is cancelled on account of the differing direction of current flow in the ring-shaped base in the area of the axis of symmetry of the antenna; however, outside of this axis the compensation of this field is incomplete and an additional magnetic field is generated in the z direction. Though this field is substantially not relevant for the generation of magnetic resonance signals, but considering the specific absorption rate (SAR), it definitely influences the total balance negatively and must also be taken into the calculation when coupling to other antennas.
A coupling to other antennas can occur in various customary antenna combinations such as are used, e.g., for “whole-body MRI” or for “continuous arterial spin labeling” (CASL). This last-named technology is used for perfusion measuring, see [7]. In it, in addition to the antenna structure detecting the brain an additional antenna structure is used as a so-called “label antenna” (conventionally a surface coil) on the patient's neck (or of the test person or test animal) in order to magnetically mark the blood at the carotid artery (by inverting the nuclear magnetization in the blood) in order to represent its subsequent local distribution in the brain by MRI. The latter takes place by means of the antenna detecting the brain.
Since the magnetic field generated from the ring-shaped base of a helmet antenna also has components vertical to the z axis, a superpositioning with the B1 field generated by the holders occurs at least in areas outside of the axis of symmetry and therefore a deterioration of the homogeneity of this B1 field occurs. Furthermore, the currents of the two holders also result in a significant stray field outside of the brain, which may amplify the coupling to other antennas (e.g., label antennas).